Method and system for nuclear imaging using multi-zone detector architecture

ABSTRACT

A method and system for nuclear imaging normally involves detection of energy by producing bursts of photons in response to interactions involving incident gamma radiation. The detector sensitivity is increased by as much as two orders of magnitude, so that some excess sensitivity can be exchanged to achieve unprecedented spatial resolution and contrast-to-noise (C/N) ratio comparable to those in CT and MRI. Misplaced pileup events due to scattered radiation are rejected for each of the central groups to reduce image blurring, thereby further improving image quality. The reduction in detector thickness minimizes depth-of-interaction (DOI) blurring as well as blurring due to Compton-scattered radiation. The spatial sampling of the detector can be further increased using fiber optic coupling to reduce effective photodetector size. Fiber-optic coupling also enables to increase the packing fraction of PMTs to 100% by effectively removing the glass walls.

CROSS-REFERENCE TO RELATED APPLICATIONS

The present continuation patent application claims priority benefit ofthe U.S. patent application serial number 12706707 entitled “A Methodand System for Nuclear Imaging Using Multi-Zone Detector Architecture”,filed on Feb. 16, 2010 under 35 USC 111(a), which in turn claimspriority benefit of the U.S. provisional application for patent serialnumber 61246918 filed on Sep. 29, 2009 under 35 U.S.C. 119(e). Thecontents of these related applications are incorporated herein byreference for all purposes to the extent that such subject matter is notinconsistent herewith or limiting hereof.

FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

Not applicable.

REFERENCE TO SEQUENCE LISTING, A TABLE, OR A COMPUTER LISTING APPENDIX

Not applicable.

COPYRIGHT NOTICE

A portion of the disclosure of this patent document contains materialthat is subject to copyright protection. The copyright owner has noobjection to the facsimile reproduction by anyone of the patent documentor patent disclosure as it appears in the Patent and Trademark Office,patent file or records, but otherwise reserves all copyright rightswhatsoever.

FIELD OF THE INVENTION

The present invention relates generally to the field of radiationimaging with emphasis on medical applications to nuclear medicine. Moreparticularly, the invention relates to a method and an apparatus forhigh resolution/sensitivity and improved image signal-to-noise ratio(SNR) for planar and tomographic (PET/SPECT) imaging.

BACKGROUND OF THE INVENTION

Medical imaging today includes radiography and computed tomography (CT)using x rays, nuclear imaging of injected or ingestedradiopharmaceuticals using scintillation cameras and SPECT or PETscanners, magnetic resonance imaging (MRI) using strong magnetic fields,and ultrasound (US) imaging using high-frequency sound waves. With theexception of nuclear imaging, all the medical imaging modalities rely onthe fact that the energy penetrating the body's tissues interacts withthose tissues and the images provide anatomical, or structural,information about the tissues. In radiography, an external sourceproduces an intense beam of x-rays, which is passed through the body,and as part of the beam is absorbed to varying degrees by differenttissues, different parts of the film are sensitized to different degreesand an image of anatomical structure is obtained. In the case of nuclearimaging with gamma rays, the radiation is emitted by sources introducedinto the body. Gamma rays differ from x-rays in that x-rays are emittedby orbital electrons in atoms, while gamma rays are emitted from withinthe nuclei of atoms and can also have considerably higher energies thanx-rays. In nuclear imaging, the radiopharmaceuticals are formed byattaching a radioactive tracer to a pharmaceutical known topreferentially accumulate in the organ of interest. Since high-energygamma radiation can penetrate bone and soft tissue alike, the pattern ofthe emerging radiation is a reflection of the distribution of theadministered radiopharmaceutical, and provides functional informationabout blood flow, metabolism, or receptor density within the organ ofinterest.

Nuclear imaging is of two types, single-photon imaging and annihilationcoincidence, or positron imaging. Single-photon imaging is used in gamma(or scintillation) cameras and in single-photon-emission computedtomography (SPECT), while annihilation coincidence imaging (ACD) is usedin positron-emission tomography (PET). Nuclear images may be planar ortomographic, planar images essentially being two-dimensional maps of theradioisotope distribution, while single photon emission computedtomography (SPECT) is the tomographic counterpart of planar nuclearimaging, and produces an image of source distribution through a sectionof the body. In either case, x-rays or gamma-rays emerging from a seriesof different angles through the body are used to reconstruct a series oftomographic images. SPECT images enable physicians to make more accurateassessment of the functional state of specific organs or tissues withinthe body. The same radioactive isotopes are used in both planar andSPECT imaging. Since gamma rays from a source will be emitted equally inall directions, a parallel-hole lead collimator is needed to prevent thephotons from reaching the detector by any path other than through theholes. The collimators typically consist of thick plates of lead withnarrow parallel holes through which the gamma rays can pass. Gamma raysnot traveling parallel to the holes are absorbed, or stopped by the leadbefore reaching the detector. Thus, if a collimator is placed over thebody, a one-to-one correspondence between interaction points in thedetector and the distribution of the isotope within the body can beestablished, which enables a planar image of this distribution to beobtained. If the collimator were to be removed the gamma rays wouldreach the detector from all directions, a uniformly white image withlittle information content would be obtained. In the case ofpositron-emission imaging, the gamma rays themselves do not comedirectly from the nucleus. The nucleus emits a positron, or positiveelectron, which has only a short half-life and annihilates with anorbital, or negative electron, resulting in the emission of a pair ofoppositely-directed 511 keV gamma rays. The fact that the pair of gammarays is always emitted along a straight line makes it unnecessary to usecollimators as in single-photon imaging. It is only necessary to use twodetectors placed on opposite sides of the positron source to determinethe line along which the photons are emitted. Image reconstructionsoftware used in PET then enables a tomographic image of the sourcedistribution to be generated. A state-of-the-art PET scanner typicallyutilizes banks of discrete stationary detectors surrounding the patient,so that annihilation photon pairs can be recorded by detector pairs fromall projection angles. Since the positions of the positron emitters liealong the lines of response (LOR) of the detector pairs, a parallel-holecollimator to limit the direction of the photons is not needed, andattenuation by the collimator is avoided. A PET scanner system istherefore more sensitive to the presence of radioisotopes than SPECTcameras, and enables more subtle pathologies to be detected.

Although nuclear imaging systems are unique in their ability to providefunctional information, in contrast to the other imaging modalities,they have the disadvantage of having the lowest spatial resolution,spatial resolution being the size of the smallest object that can beresolved in the image. While MRI and CT imaging can typically provideresolutions of 1.0 and 0.4 mm, respectively, planar nuclear cameras orSPECT scanners provide resolutions of about 7 mm, while PET scannersprovide a resolution of 4-6 mm. The lower spatial resolution of nuclearimages, coupled with their inability to provide anatomical information,has led to the development of hybrid PET/CT or PET/MRI systems in asingle gantry, with a single bed for the patient, making co-registrationof the PET images with those of CT or MRI possible. The anatomicalimages from CT or MRI provide more accurate information regarding thelocations of lesions in the organs of patients, CT images additionallyalso providing data for the attenuation correction of the PET data. Itis evident that for the quality of nuclear images to be more competitivewith those of CT and MRI, the detector systems for them requiresignificant improvement in spatial resolution, sensitivity and imagesignal-to-noise (SNR).

Scintillation detectors emit visible light photons when gamma radiationis detected. Nuclear images are built up by counting the number of gammaray interactions for each pixel in the image, and image contrast arisesfrom differences of count density in the image. Detection of a gamma rayinteraction in a scintillator consists in converting the burst ofvisible light photons into electrical pulses using a photodetector, suchas the photomultiplier tube (PMT) or a photodiode (PD), each providedwith a preamplifier (PA). The total energy released by a gamma ray isobtained by adding all the PA outputs of exposed PMTs in a summingamplifier (SA), whose output is the energy signal. Since the amplitudeof the PA outputs is highest for those PMTs closest to the interactionpoint and decreases with distance, the spatial coordinates of this pointare determined by using the PA outputs for locating the centroid oflight absorption. The circuit used to generates the (x,y) coordinates,or position signals, is referred to as Anger position logic. A gamma raycan interact with the scintillator by depositing its full energy at onceby photoelectric interaction, or in smaller fractions by Comptonscattering. Photoelectric interaction is analogous to a fast-movingbilliard ball hitting a stationary ball and being stopped, andtransferring all its kinetic energy to the stationary ball. In the caseof a gamma ray, the energy is transferred to an orbital electron, whichthen immediately releases its energy in the form of light in a series ofcollisions with atoms within a short distance. For all practicalpurposes, the light can be considered to be emitted from the interactionpoint of the gamma ray. In photoelectric interaction the number ofvisible light photons, and hence the amplitude of the energy signal atthe SA outputs is a maximum. Compton-scatter interaction, on the otherhand, is analogous to the fast-moving billiard ball colliding with thestationary ball at a glancing angle, thus losing only part of its energyand changing its direction. The scattered gamma ray may interact withthe scintillator at any distance from the initial point of interaction,or even escape from the scintillator crystal undetected. The energysignals of Compton-scatter events lie anywhere between a maximum and aminimum. In practice, nuclear medicine images are created using thoseenergy signals that are within 10% to 20% of the most probable energysignal magnitude, which is not necessarily the same as that of aphotoelectric event. Valid energy signals are selected by applying theSA outputs to a pulse-height analyzer (PHA), which accepts only energysignals within the energy window chosen.

Anger type scintillation cameras for planar imaging consist in a flatlarge-area detector viewed by an array of PMTs and associatedelectronics for the determination of the gamma-ray energy and thespatial coordinates of interaction points. Anger cameras are photoncounting systems and operate in pulse mode, so that images are acquiredone interaction at a time. This is in contrast to current-mode operationin the other modalities, in which images are acquired as part of asingle operation. In the basic scintillation camera, only a single SA isprovided, with the inputs being the PA outputs of all the PMTs in thedetector, typically 37, 61, or 91 in number. A single SA for the wholedetector means that only one gamma ray interaction at a time can bedetected, and detector operation is in single-zone mode. Energydetermination is achieved by summing the outputs of all the PMTstogether, while the spatial coordinates are determined by generatingfour position signals, usually referred to as the X+, X−, Y+, and Y−signals. The X+ and X− signals are obtained by summing the outputs ofthe PMTs in the right and left halves of the PMT array, respectively,while the Y+ and Y− signals are obtained by similarly summing theoutputs of the upper and lower halves of the PMT array, respectively.The X and Y coordinates of each event are then obtained takingX=(X+)−(X−) and Y=(Y+)−(Y−), normalized to (X+)+(X−) and (Y+)+(Y−),respectively. It is therefore inevitable that the determination of thephoton energy and the spatial coordinates becomes susceptible to noisefrom PMTs at large distances from the interaction point. Furthermore,the fact that each event detected involves the entire array of PMTs inthe detector, and that only one interaction at a time can be detected,remains one of the major shortcomings of the conventional scintillationcamera. Two or more interactions in the detector at the same time leadto energy signal that are too large and fall outside of the energywindow, and are therefore rejected. A source of image blurring is whentwo or more Compton-scatter interactions at different points in thedetector occur simultaneously and the energy signals add up to thatresulting from photoelectric interaction. The composite energy signalwould then be accepted by the PHA as valid, while the Anger positioninglogic circuitry generates the coordinates of an intermediate location asbeing the point of a photoelectric interaction. These events arereferred to as misplaced pileup events, since a photoelectricinteraction has not occurred there, the effect of these events being toblur the image of the radiopharmaceutical distribution. Considerableeffort and research has been devoted to increasing the number ofgamma-ray interactions that can be detected at a time. One solutionadopted in state-of the-art nuclear cameras is to dispense with the SA,digitize each of the PA output of the PMTs individually in separateanalog-to-digital converters (ADCs) and read the data into computermemory. A computer program then analyzes the data to identify groups ofPA outputs that correspond to valid gamma-ray interactions. Thistechnique has helped in that up to three events at a time can now bedetected, and cameras that operate in this way are referred to asdigital cameras, which are normally advertised as having one ADC per PMTto reflect this improvement in performance. Another approach to gammacamera design that overcomes the limitations of single-zone operation isone in which the detector is divided into multiple geographical zones,each of which is provided with an SA and Anger position logic, andoperates independently, so that multiple events can be detected. Thisapproach, however, is also limited to detecting at most only three validevents at a time. The large number of valid interactions that continuesto be rejected, therefore, remains a significant disadvantage.

Detector crystals for PET are considerably thicker than forsingle-photon imaging due to the higher energy of positron annihilationgamma rays. Additionally, the fact that collimators are not used in PET,and that the gamma rays can also reach the detectors at oblique anglesmeans that the greater the angle of incidence, the greater the distancesthat can be traversed by the gamma rays in the crystal. Imagereconstruction software in PET assumes that the scintillation eventoccurs directly below the point of initial incidence on the crystalsurface, which is very rarely the case. Mispositioning of scintillationevents in this way leads to image blurring referred to asdepth-of-interaction (DOI) error. An additional source of error is thatfull-energy Compton-scatter interactions, which are also used in imageformation, are not necessarily located along the initial trajectory ofthe gamma rays and contribute to the blurring. Thus, thick detectors forPET have the attendant problems of lower spatial resolution andsignificant blurring of the image. The resolution of PET images iscurrently better than those of planar images because PET detectormodules as a rule are divided into small segments typically 3×3×30 mm³to improve spatial resolution by limiting the divergence of thescintillation light once an interaction in a segment has taken place.Detector segmentation, however, typically has no effect on blurring dueDOI error and Compton scattering since the high energy gamma rays areable to penetrate the segments regardless.

Another important feature of nuclear imaging systems is count-ratecapability, which is directly related to the detector's deadtime.Imaging applications can involve patient movement or fast redistributionof the radiopharmaceutical within the body, which can lead to blurredimages unless an image can be built up in the shortest possible time.The shortest interval between individual detected events in thescintillation camera is referred to as its deadtime. The deadtime coversthe period between the gamma-ray interaction in the crystal and thetransfer of digitized energy and positional information to computermemory. There are two types of deadtime, the paralyzable deadtime of thedetector, which is a characteristic of the scintillator decay time, andnonparalyzable deadtime, which is the time needed by thesignal-processing hardware and the computer interface to generate andtransfer the digital data into computer memory. Nonparalyzable deadtimeis generally longer, but remains constant, while paralyzable deadtimeincreases with count-rate. The count rate for a paralyzable detectortherefore increases only up to a peak value, which is proportional tothe reciprocal of its deadtime τ₀, and then decreases as count rateincreases, eventually leading to detector paralysis. This occursbecause, when one scintillation event follows immediately after theprevious one, they merge together and the energy signals pile up in theSA, leading to the rejection of both by the PHA. Pile-up rejectionbecomes more frequent and the deadtime longer as the count rateincreases. The way in which detector paralysis is prevented in existingimaging systems is by limiting the count rate, and consequently also thedose of administered radiopharmaceutical. The count-rate is rarelyincreased above the 20% count loss point, which means that only 80% ofthe peak count rate can be attained. The inefficient use of the detectorand PMT array has long been recognized and has been receivingconsiderable attention. However, as noted earlier, it has not beenpractical to increase detector sensitivity by more than a factor of twoor three times, and only two or three events per deadtime period can berecorded.

PET imaging requires that the interaction times of events in acoincidence pair be compared in order to establish a temporal overlapbefore use in image reconstruction. Current designs compare theinteraction times of individual events against a master clock and storethe time stamps along with the energy and spatial coordinates, forcoincident events to be identified subsequently by software. The timewindow for coincidences is significantly shorter than the detectordeadtime, as a result of which only a tiny fraction of detected eventswill lead to coincidences. This means that rapid redistribution ofradioactivity within the body, which may take place in manyinvestigations, cannot be observed in real time, but only after thefact, since the data typically must first be processed using imagereconstruction software.

Scintillation detectors for 140 keV gamma rays normally have thicknessesof 6-12 mm in order to ensure adequate sensitivity. Intrinsic spatialresolution and detector sensitivity, however, have conflicting thicknessrequirements, since the greater the detector's thickness, the greaterthe divergence of the scintillation light that emerges on the PMT side.Detector sensitivity needs to be enhanced in other ways if intrinsicresolution is not to be degraded. A 7.5-mm thick NaI detector for 140keV gamma rays will have an intrinsic resolution of 7.2 mm for planarimaging and SPECT. BGO detectors for 511-keV gamma rays typically have20-30 mm thicknesses to have adequate detection efficiency, and acontinuous BGO detector of 30 mm thickness would have an intrinsicresolution of 31 mm, which is unacceptable. State-of-the-art PETscanners therefore achieve 4-6 mm spatial resolution by employing2.54×2.54 cm² blocks of BGO or LSO segmented into 8×8 arrays of (3×3mm²) elements by means of saw-cuts. The saw-cuts between segments aresilvered to help prevent divergence of the scintillation light as itemerges. Further improvement of spatial resolution would require greaterreduction of the segment width, which makes it progressively moredifficult to identify individual segments if shared photodetectors arebeing used. Experts in the field had estimated that 2-mm isotropicresolution using this method would be the limit.

Although detector segmentation is effective in restricting divergence ofthe scintillation light once an interaction in a segment has takenplace, it has no effect on the ability of obliquely incident gamma raysand Compton-scattered interactions to penetrate the detector segments.Image distortion due to gamma rays penetrating the detector segments isreferred to as depth-of-interaction (DOI) error. Correction of DOIerrors requires detector technology that involves either multiple layersof scintillator or the use of avalanche photodiodes (APDs) asphotodetectors along with the PMTs, which increases both detector costand complexity without improving resolution. Although positrontomographs with correction for DOI errors are becoming available,improvement in image quality has only been marginal. Tests onhigh-resolution prototypes have also shown that a limit to theresolution obtainable is set by what is referred to as the “blockeffect”, although later investigation has shown that that this effect isnot limited to just block detectors but also extends to discretesegments. Current strategies to overcome the limitations of nuclearimaging systems can be regarded as treating the symptoms, instead ofcuring the curing the disease itself Modern PET scanners are also beingprovided with TOF capability to limit image noise arising from thebackprojection of data for image reconstruction. Image quality innuclear imaging, nevertheless, remains significantly inferior to thoseof CT and MRI. Currently available detector architecture is unlikely toaddress the problem to any significant extent. Moreover, despite thelarge numbers of publications and patents awarded for the improvement ofsensitivity in nuclear imaging systems, little attention has beendevoted to improving spatial resolution. It is also to be noted thatidentification of valid coincidences in current PET systems is carriedout using software operating on large amounts of raw data as opposed toreal-time validation using hardware, which leads to an inefficient useof storage space and computing time.

A direct approach to achieving multi-zone operation has been thedivision of the detector into multiple independent geographical zones,as is the case with pixellated arrays of NaI(Tl) and CsI(Tl)scintillators, for instance. Discrete detectors in these arrays,however, require independent readout using Si photodiodes, which makesthem expensive due to their large numbers. Such arrays find applicationin compact, mobile cameras and pulse-height spectrometry systems.Semiconductor detector arrays of high purity germanium (HPGe) andZnCdTe, which eliminate the need for a scintillator-photodetectorcombination, have also been developed with varying degrees of success.HPGe detectors require cryogenic cooling, however, while ZnCdTedetectors are currently too difficult to fabricate in large size. As aresult, despite its low quantum efficiency, unit-to-unit variations inquantum efficiency, and poor packing fraction from the dead zone due tothe glass walls, the PMT remains the photodetector of choice in nuclearimaging systems.

In view of the foregoing, there is a definite need for a new approach todetector design to improve sensitivity, increase image SNR and spatialresolution. Lesion detectability in images needs to be further improvedby reducing/eliminating blurring due misplaced pileup events in planarand tomographic imaging, and due to DOI error and Compton-scatterinteractions in PET systems. Theoretical investigation of factorsaffecting resolution shows that a fundamental limit to resolution incontinuous as well as segmented detectors that worsens with detectorthickness and segment length, respectively, is set by Compton-scatterinteractions, and that segmentation becomes progressively less effectiveas resolution increases. Image granularity due to the dead space betweenthe segments also worsens as segment size decreases. The need forexchanging detector sensitivity for increased resolution suggests thatthe use of a continuous detector with multizone architecture in PETsystems will be less complex and also more cost-effective.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention is illustrated by way of example, and not by wayof limitation, in the figures of the accompanying drawings and in whichlike reference numerals refer to similar elements and in which:

FIGS. 1A-1E illustrate exemplary event distributions and the detectorzones within PMT array that are activated to detect them in accordancewith embodiments of the present invention;

FIG. 2A illustrates a simplified block diagram of exemplary misplacedpile-up suppression circuitry for detector zones in accordance with anembodiment of the present invention;

FIG. 2B illustrates a simplified block diagram of exemplarydetector-zone circuitry in accordance with an embodiment of the presentinvention;

FIG. 3 illustrates a simplified block diagram of an exemplaryscintillation camera with MZDA in accordance with an embodiment of thepresent invention;

FIG. 4A illustrates a maximum usable field-of-view (FOV) for a hexagonalPET scanner in accordance with an embodiment of the present invention;

FIG. 4B illustrates a minimum circular field-of-view (FOV) for ahexagonal PET scanner in accordance with an embodiment of the presentinvention;

FIG. 4C illustrates a minimum circular field-of-view (FOV) for ahexagonal PET scanner based on curved detectors in accordance with anembodiment of the present invention;

FIG. 5 illustrates exemplary coincidence validation logic circuitry fora PET/SPECT scanner in accordance with an embodiment of the presentinvention;

FIG. 6 illustrates a simplified functional block diagram of an exemplaryhexagonal PET/SPECT scanner in accordance with an embodiment of thepresent invention;

FIG. 7A illustrates exemplary geometric sampling employing a largernumber of reduced effective-size PMTs by using fiber-optic coupling inaccordance with an embodiment of the present invention;

FIG. 7B illustrates an exemplary equivalent circuit of PMT P₁ and its PAin FIG. 7A in accordance with an embodiment of the present invention;

FIG. 7C illustrates an exemplary equivalent circuit of a four-channelPMT, associated PAs and SA in accordance with an embodiment of thepresent invention; and

FIG. 8 illustrates an exemplary circuit enabling four adjacent radiationcounters to share an amplifier, PHA and ADC in accordance with anembodiment of the present invention.

Unless otherwise indicated illustrations in the figures are notnecessarily drawn to scale.

SUMMARY OF THE INVENTION

To achieve the forgoing and other objectives and in accordance with thepurpose of the invention, a method and system for nuclear imaging usinga floating-boundary, multi-zone detector architecture (MZDA) ispresented.

In one embodiment of MZDA a system includes at least one energy detectorfor producing bursts of optical photons in response to events includingincident rays of radiation. MZDA includes F number of sharing groups ofseven photodetectors arranged in a honeycomb array for viewing zones ofthe at least one energy detector and converting the bursts of photonsinto signal outputs, where each of the central groups is associated witha zone of the at least one energy detector. A misplaced pile-upsuppression circuitry is associated with each of the central groups forrejecting signal outputs that are due to scattered radiation of theincident rays. A detector-zone circuitry is associated with each of theF number of sharing groups for generating energy and position signals indetection of F number of the events during the period of one detectordeadtime. Another embodiment further includes pileup preventioncircuitry, which enables the paralyzable detector to operate asnonparalyzable, and prevents the deadtime from increasing with countrate and resulting in detector paralysis. Each of the above embodimentsincluded associated analog-to-digital convertors (ADCs) for the energyand position signals. A high-speed multiplexer with F inputs transfersthe digital energy and position signals of the F number of the events tocomputer memory. Still another embodiment further includes spatiallinearity correction hardware, which operates on the digitized signalsto mitigate distortions arising from nonlinearities. Another embodimentfurther includes energy correction hardware for mitigating errorsarising from inequalities in gains of the photodetectors for imagedisplay and data analysis. The above techniques of MZDA to improve thesensitivity, spatial resolution, and signal-to-noise ratio (SNR) ofnuclear images have been applied to exemplary designs for both planarand tomographic (PET/SPECT) imaging.

In another embodiment of MZDA a method includes steps for using fiberoptics to couple the bursts of photons from a continuous scintillator tophotodetectors for reducing the effective size of the photodetectors,thereby also effectively removing the glass walls from PMTs andincreasing the packing fraction to 100%, enabling to substantiallyincrease the number of photodetectors and further reduce effectivedetector deadtime. This increases both the detectorsensitivity/count-rate response and improves the spatial sampling of thedetector, hence enabling more accurate determination of the energy andposition signals and reducing image blurring due to noise. Anotherembodiment further includes steps for coupling independent adjacentradiation detectors or particle counters to photodetectors. This enablesadjacent radiation detectors or particle counters, normally providedwith separate signal processing circuitry, to be arranged into detectorzones and to share the circuitry for cost reduction while maintainingthe spatial sampling capabilities of the independent circuitry.

Other features, advantages, and objectives of the present invention willbecome more apparent and be more readily understood from the followingdetailed description, which should be read in conjunction with theaccompanying drawings.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

The present invention is best understood by reference to the detailedfigures and description set forth herein.

Embodiments of the invention are discussed below with reference to theFigures. However, those skilled in the art will readily appreciate thatthe detailed description given herein with respect to these figures isfor explanatory purposes as the invention extends beyond these limitedembodiments. For example, it should be appreciated that those skilled inthe art will, in light of the teachings of the present invention,recognize a multiplicity of alternate and suitable approaches, dependingupon the needs of the particular application, to implement thefunctionality of any given detail described herein, beyond theparticular implementation choices in the following embodiments describedand shown. That is, there are numerous modifications and variations ofthe invention that are too numerous to be listed but that all fit withinthe scope of the invention. Also, singular words should be read asplural and vice versa and masculine as feminine and vice versa, whereappropriate, and alternative embodiments do not necessarily imply thatthe two are mutually exclusive.

The present invention will now be described in detail with reference toembodiments thereof as illustrated in the accompanying drawings.

Detailed descriptions of the preferred embodiments are provided herein.It is to be understood, however, that the present invention may beembodied in various forms. Therefore, specific details disclosed hereinare not to be interpreted as limiting, but rather as a basis for theclaims and as a representative basis for teaching one skilled in the artto employ the present invention in virtually any appropriately detailedsystem, structure or manner.

It is to be understood that any exact measurements/dimensions orparticular construction materials indicated herein are solely providedas examples of suitable configurations and are not intended to belimiting in any way. Depending on the needs of the particularapplication, those skilled in the art will readily recognize, in lightof the following teachings, a multiplicity of suitable alternativeimplementation details.

The three broad design objectives of the preferred embodiment of thepresent invention are (a) the improvement of detector sensitivity toincrease image SNR, (b) the improvement of spatial resolution toincrease the detectability of lesions in nuclear images, and (c) thereduction of image blurring due to misplaced pileup events inscintillation camera and SPECT images and to DOI error andCompton-scatter interactions in PET images. This is achieved bysignificantly reducing detector deadtime using multi-zone detectorarchitecture (MZDA) and including misplaced pileup suppression (MPS)circuitry. Whereas the basic nuclear imaging system builds up an image asingle scintillation event per deadtime period, MZDA reduces effectivedetector deadtime by a factor S, which can have as much as two orders ofmagnitude (S≧100), thereby increasing the peak count rate by a factor ofS times. In preferred embodiments of the present invention, using ahoneycomb arrangement preferably of 7 PMTs as a detector zone, S can beestimated as S≈P/11, where P is the PMT array size. The resultingincrease in detector sensitivity obviates the need for thick detectorsto improve sensitivity, thereby enabling the spatial resolution to bealso improved. Preferred embodiments with MZDA and MPS circuitrygenerally reduce misplaced events in the image by a factor of P times.In the preferred embodiments, detector sensitivity, which can be made ashigh as S times if desired, is limited to a factor just two to threetimes, so that excess sensitivity may be exchanged for improved spatialresolution. The extent, to which S can be increased, in preferredembodiments using MZDA, depends on the minimum PMT size available, whichis currently 1.27 cm (½″). A scintillation camera design, in accordancewith the present invention, using a 1.7-cm thick, hexagonal 29-cm wide,NaI(Tl) detector viewed by 331 1.27-cm PMTs will provide a resolution ofabout 2.1 cm at the collimator face and 3.4 cm at 10 cm from thecollimator face, with two times the sensitivity of a single-zone designusing the same array size. Designs based on the prior art typically usecrystal thicknesses of 0.95 cm to achieve adequate sensitivity, and canprovide a resolution of at best 7 cm. Further improvement in sensitivityand spatial resolution is possible only by reducing effective PMT size,since reduction of actual PMT size cannot continue without limit becauseof the glass walls of the PMTs. In preferred embodiments of the presentinvention, effective PMT size is reduced by using fiber-optic couplingbetween the scintillator and the PMTs. In a non-limiting example, adesign using a 36×32 cm²NaI(Tl) crystal in which the effective size ofan array of P=1930 1.27-cm PMTs is reduced to 7.6 mm by means offiber-optic coupling, leads to a resolution of 1.0 cm at the collimatorface and 1.8 mm at 10 cm from the collimator face. In the presentinvention, fiber optic coupling enables the packing fraction of the PMTsto be increased to practically 100% since the thickness of the glasswalls no longer constitutes a problem. An additional benefit of fiberoptic coupling, in the present invention, is therefore that PMTs can bechosen on the basis of performance and cost, irrespective of size andshape.

In accordance with the present invention, a high-resolution/sensitivityhexagonal PET/SPECT scanner using continuous-crystal NaI(Tl) detectorswith MZDA, MPS and pileup-prevention circuitry (PPC) shows that thescanner's sensitivity can further be increased by operating threedetector pairs simultaneously to triple the coincidence detection ratein PET and to achieve six times the count-rate capability of a singlerotating detector head in SPECT imaging. In current PET designs, thespatial (x-, y-, z-) coordinates of valid gamma-ray interactions arefirst stored in computer memory along with time stamps using a masterclock, and the raw data post-processed using software to extract thesmall fraction of valid coincidences in opposing detector pairs. Asignificant feature of the present invention is that valid coincidenceswithin the field-of-view (FOV) are identified and time-of-flight (TOF)information determined in real time using hardware, which savesdata-processing time. In the present invention, coincidence validationinvolves pulses of a fixed duration τ_(c) being triggered by each validgamma-ray interaction, and then being compared in real time by thedigital coincidence detection circuitry with pulses from opposingdetector zones within the FOV to test for valid coincidences. Thisenables coincidences in each of the three detector pairs to be validatedsimultaneously, thereby increasing the coincidence detection rate of asingle detector pair with MZDA by a factor of three times. In thepresent invention, coincidence events are accepted as valid only if theindividual lines of interaction (LOIs) between opposing detector pairsare within the angular FOVs of the detector zones. The angular FOVremains in the range 29° to 34°, as shown in FIG. 4, for a patient ofdiameter D=30 cm and detectors of length L≈2D≧60 cm in a non-limitingexample. Restriction of the LOIs to within the angular FOV, in additionto employing NaI(Tl) to improve energy resolution, minimizes imagedegradation due to accidental and scatter coincidences from outside theFOV. In a non-limiting example, using a 0.8-mm thick scintillator and anarray of P=1287, 1.27-cm PMTs per detector provides a spatial resolutionof 0.66 mm and a relative sensitivity of better than 2 times that ofsingle-zone operation, or a resolution of 1.5 mm and a relativesensitivity of better than 10 times, if the crystal thickness isincreased to 1.8 mm. In PET, block detector thicknesses of 2-3 cm leadto large DOI errors. In the present invention, image blurring from thissource is virtually eliminated on account of the significantly reducedscintillator thicknesses, so that correction of DOI errors is no longerneeded. The second major cause of image noise in PET systems is thecontribution of random and accidental coincidences to image data. In anon-limiting example, a PET design uses a hexagonal arrangement ofcontinuous flat NaI(Tl) crystals as the detectors. Since NaI(Tl) hassignificantly higher light output than most other scintillators, thisenables an energy window of ±10% to be employed, as compared to ±20% forBGO, for instance. The reduced energy window leads to substantialreduction in the scatter and random coincidences added to the image andtherefore to an increase in the SNR. In other embodiments of the presentinvention, the NaI(Tl) detectors can be replaced with those of asuitably fast scintillator, such as, but not limited to, LSO, LYSO, GSO,etc., if time-of-flight (TOF, or TF) capability is desired.

In a preferred embodiment of the present invention, a PET scanner basedon the 0.8-mm thick NaI(Tl) scintillators can also be employed as aSPECT scanner with the addition of matching low-energy high-resolutioncollimators. In non-limiting example, the six detectors are designed tooperate independently alongside of each other. The resolution of thescanner is reduced to 0.95 mm at the collimator face and to 1.6 mm at 10cm from the collimator face, the relative sensitivity being 6.0 timesthat of a single rotating detector head operating in single-zone mode. Adecrease of collimator thickness to get a reduced resolution of 2.0 mmat 10 cm from the collimator face improves the relative sensitivity to17, as compared to just the 6.0 times that of a single detector head.The SPECT scanner can additionally be employed for planar imaging ifheld stationary during imaging. In this embodiment, each pair ofopposing detector heads can be operated back-to-back to acquire threehigh quality images oriented at angles of 60° to one another. Spatialresolution remains the same as for SPECT imaging, while thesensitivities of the three detector pairs will be one-third that forSPECT operation. In prior art designs, resolution in the planar imagesnormally decreases progressively on moving away from the collimator. Inthe present invention with two opposing detectors operatingback-to-back, the resolution of planar images decreases only on goingtoward the scanner axis, which leads to significantly better imagequality.

In various embodiments of the present invention, the implementation ofMZDA need not be limited to single-crystal detectors as discussed above.Detector zones in continuous-crystal designs are needed because of thelarge size and poor spatial resolution of PMTs used as thephotodetectors. Detector zones permit the spatial coordinates ofscintillation events to be determined using Anger positioning logic.Pixellated scintillator arrays for compact mobile cameras andsemiconductor detector arrays for energy-selective radiation countingare of sufficiently small size to provide high resolution, and aretherefore provided with their own dedicated amplifiers and ADCs, whichmakes them particularly expensive. Counter systems in which thecount-rate density is much less than the pixel density can be redesignedto benefit from MZDA, in accordance with the present invention, in whichgroups of the radiation or particle detectors can sharesignal-processing circuitry, thereby offering a more economical designthat retains the original spatial resolution. In a non-limiting example,the detector density is N₀ (pixels/cm²) and the expected maximumcount-rate density for the application is C₀ (counts/s/cm²), whereC₀<N₀. The true count rate C₀ can be recorded with no more than 20%deadtime losses, provided that C₀≧0.223/τ₀, where τ₀ is the deadtime ofa single detector in the pixel matrix operating independently, whichmeans that S=4.48C₀. Amplifiers and event-validation circuitry in apixellated detector can then be shared between two or more adjacentpixels provided that k=N₀/S≧2, or N_(o)≧2S=8.96C_(o)≈9.0C_(o). If 2≦k<3,two adjacent pixels can share the same circuitry, since at most only oneevent on average arrives at each pixel pair. Similarly, if 3≦k<4, threeadjacent pixels can share the circuitry, and so on. The electricalsignals generated in the shared signal-processing circuitry are easilydecoded for event position within the pixel matrix, while the energy ofthe gamma rays is determined from the digital values of these signals.The ability to share the circuitry between multiple pixels enables thepixels to be made sufficiently small for good spatial resolution whilereducing the circuit cost by a factor of k times.

FIGS. 1A-1E illustrate exemplary PMT array structures in accordance withembodiments of the present invention. FIG. 1A depicts an exemplary PMTarray 100 with a honeycomb structure to make detector zones that consistof groups of seven PMTs. The seven PMTs 1-7 provide sufficientinformation for accurate energy and position signals for the detectedevent to be generated. Analysis shows that, when a valid event isdetected at the central PMT 1, the seven closest PMTs collect between90% and 96% of the scintillation light, and that cross-talk between twoadjacent detector zones is less than 1%. FIG. 1B depicts the lowestmaximum number of 7 independent detector zones for the array 110 of 132PMTs shown, while FIG. 1C shows the highest maximum number of 15 ofindependent detector zones in array 120, depending on eventdistribution. The number of independent detector zones is a maximum whenno new ones can be added without causing an overlap. FIG. 1D illustratesthat detector zones can be tightly packed together in array 130, likezones 1 to 6, completely isolated, like zone 7, or overlappingspatially, like zones 8 and 9, which will be rejected by thedetector-zone circuitry. FIG. 1E illustrates PMTs around the outside ofthe array 140 cannot become central PMTs to detector zones because therewill then be fewer than seven PMTs in the detector zones. Possiblecentral PMTs to detector zones can be grouped together to share detectorzone circuitry since no two PMTs separated by less than two interveningPMTs can be central PMTs to independent detector zones. The PMTs of adetector zone obviously meet this criterion, and the maximum number ofcentral PMTs that can share detector-zone circuitry is therefore alsoseven. The number F of detector zone circuitry needed can be minimizedby maximizing the number of sharing groups consisting of seven centralPMTs. Circuits for groups of central PMTs sharing detector zones cancome in modular form with the modules consisting of groups 2 to 7 PMTs.Any PMT array can be formed by joining F such modules together. The F=19groups of central PMTs for the array in this example are made up ofmodules consisting of 2 to 7 PMTs, although a module with 6 PMTs was notneeded.

FIG. 2A illustrates a simplified block diagram of exemplary misplacedpile-up suppression circuitry in accordance with an embodiment of thepresent invention.

FIG. 2A shows misplaced pile-up suppression (MPS) circuitry 210 andevent switches 230 for a group of 7 central PMTs sharing detector-zonecircuitry, since the detector-zone circuitry ensures that just onecentral PMT at a time is active. The function of the MPS circuitry 210is to reject preamplifier (PA) outputs of the central PMT that are dueto scattered radiation. The PA outputs of the 12 PMTs closest to thedetector zone are also checked to prevent pileup events and to minimizecross-talk between adjacent detector zones. FIG. 2A shows oneimplementation of the MPS circuitry 210 for PMTs T₁ through T₇, whichcomprise the first group of central PMTs for sharing detector-zonecircuitries. The spatial coordinates (x_(0,1), y_(0,1)) of T₁ withrespect to the crystal center and the PA outputs (v_(1,1), . . . ,v_(1,7)) of the PMTs within the detector zone are taken to the commonbus 240 through the set of analog switches 230. The associated delaylines 220 allow time for Z′_(L1) to be generated and the switches to beactivated. The MPS circuitry 210 for T₁ generates the logic-high pulseZ_(1,1) when v_(1,1)>f_(min)V_(L), which also means thatv_(1,1)>v_(1,i), i=2, 3, . . . , 7 and that the detected event may bevalid. The event at T₁ is determined to be spatially independent if noevent in the 12 second-ring PMTs at the same time satisfies thecondition v_(1,i)>f_(min)V_(L), i=8, 9, . . . , 19, which leads to bothbeing rejected. The closest that a second valid event can be to that atT₁ is in one of the 18 third-ring PMTs, as for the central PMTs of thedetector zones 1, 2 and 3 in FIG. 1D. When the PA outputs of thesecond-ring PMTs meet the above requirement, D_(1,1)=0 and Z′_(L1,1) =0,so that Z′_(L1,1) closes the analog switches 230 and also appears asZ′_(L1) in the common bus 240. Other PMTs in the same group as T₁ willnot be connected to this bus, since only one central PMT at a time canbe present.

FIG. 2B illustrates a simplified block diagram of exemplarydetector-zone circuitry in accordance with an embodiment of the presentinvention. The detector-zone circuitry includes a summing amplifier (SA)250, Anger position circuitry 260, pileup prevention circuitry (PPC)270, a pulse-height analyzer (PHA) 280 with a wider than standard energywindow, a set of 4 ADCs with T/H 290 and external data latches 295. Awider than standard energy window is used here to make subsequent energycorrection possible as indicated in FIG. 3. The PA outputs of the PMTsin the detector zone, the coordinates (x_(0,1),y_(0,1)) of the center ofT₁, and the digital signal Z′_(L1) are available on the common bus 240at the top of the figure. The analog position circuitry 260 generatesthe coordinates Δx and Δy of the interaction point with respect to thecenter of T₁, so that the absolute coordinates x′₁=x_(0,1)+Δx andy′₁=y_(0,1)+Δy of the interaction point with respect to the crystalcenter can be generated. The energy signal E′₁ is likewise generated inthe SA 250 from the PA outputs within the detector zone. The analogsignals generated are applied to the PPC 270 to recover the first pulsefrom multiple-pileup interactions. The corrected energy signal E₁ istested to determine whether or not V_(L)<E₁<V_(H), in which case alogic-low pulse Z₁ will be generated that initiates digitization at theADCs 290. The digital outputs of the ADCs are held in external datalatches 295 and read sequentially into the computer interface through anF-input multiplexer.

FIG. 3 illustrates a simplified block diagram of an exemplaryscintillation camera with MZDA in accordance with an embodiment of thepresent invention. The P PMTs of the array 310 are divided into F nearlyequal groups to share F detector-zone circuitries 330. Up to Findependent events can therefore be detected, the digital values of theenergy and position signals then being held in data latches 295, FIG. 2,until read into computer memory through a high-speed multiplexer 340.The multiplexer reads the F number of data latches once every deadtimeperiod to ensure that the computer interface keeps in step with themulti-zone detector. At the top is the array 310 of P PMTs, eachprovided with an MPS circuitry 320. The PMTs in the array are arrangedinto F groups of 5 to 7 central PMTs to share detector-zone circuitries.The PA outputs of those central PMTs for which v₁>f_(min)V_(L) areapplied to the detector-zone circuitries 330 to generate the energy andposition signals for the detected events. These signals are then appliedto the PPC 270 to recover the first event in pileup interactions thatwould otherwise be rejected by the PHA 280, the PPC thus helping toreduce detector deadtime. The energy signals are then applied to thePHAs 280 for energy discrimination, after which the analog signals fromthe detector zones are applied to their respective ADCs 290 fordigitization. The digital data from the ADCs are applied to the computerinterface through an F-input multiplexer 340. Spatial linearity andenergy correction 350 are then applied to the data before transfer tocomputer memory for image display and data analysis. Spatial linearitycorrection in the image is needed to remove distortion arising fromnonlinearities in the analog positioning circuitry, inequalities in thegains of the PAs of PMTs within the detector zones, and fromimperfections in the detector crystal and the collimator. Linearitycorrection involves assigning a correction term for the positioncoordinates of each pixel in the image. Likewise, energy correction inthe image is needed to minimize errors in the energy signals arisingfrom inequalities in the gains of the PAs in the detector zones. Ascorrection involves either an increase or decrease in E₁ for each event,most energy signals within the provisional window of the PHA 280 will beretained, while some at both ends of the window will be discarded inaccordance with the lookup table. Errors in event energy are minimizedby maintaining a regular quality control procedure that ensures that thegains of the PAs of PMTs remain as nearly equal as possible.

FIG. 4A illustrates a maximum usable field-of-view (FOV) for a PETscanner employing six flat rectangular NaI(Tl) detectors in accordancewith an embodiment of the present invention. The maximum FOV 410 is thehexagonal area viewed by the six detectors A-F. FIG. 4B illustrates aminimum circular field-of-view (FOV) for a PET scanner in accordancewith an embodiment of the present invention. For a minimum circular FOV420 of diameter D=L/2 (shaded), the angular aperture for a PMT variesfrom 28.9° at the corners to 33.6° at the centers of detectors A-F. FIG.4C illustrates a minimum circular field-of-view (FOV) for a PET scannerwith six curved detectors in accordance with an embodiment of thepresent invention. For a minimum circular FOV 430 of diameter D=L/2(shaded), the angular aperture for a PMT is uniformly equal to 30° forall PMTs in the six detectors. For an FOV of diameter D=L/2, each PMT inone detector can be in coincidence with the P PMTs in within its angularFOV in opposing detectors. Limiting coincidence detection to within theminimum FOV minimizes image noise due to scatter and randomcoincidences.

FIG. 5 illustrates exemplary coincidence validation logic circuitry fora PET/SPECT scanner in accordance with an embodiment of the presentinvention. Each detector zone in one detector can have coincidences withF other detector zones within its angular FOV. The logic level signal Zand the analog signals E, x and y are taken from the PHA 280 and theenergy and position circuitry 260, respectively, as shown in FIG. 2B.The signals labeled with the subscripts 1 i and 2 j (1≦i≦F, 1≦j≦F)correspond to detector zone i in detector 1 and to detector zone j indetector 2, respectively. Z-signals from valid interactions in eitherdetector trigger the generation of narrow pulses of duration τ_(c) 510where 2τ_(c) is the resolution, or coincidence time window, of thescanner. The τ_(c) pulse due to, say, event U from detector zone 1 i iscompared in the logic circuitry block (LC_(2j)) 525 with the τ_(c) pulse510 of every detector zone 2 j, and if a valid coincidence with just oneevent, say V, in detector zone 2 j is detected, a logic-high pulse ispresented at the corresponding output terminal in the LC_(2j) block 525.This logic-high pulse triggers a pulse of duration T_(p) 530, which isthen used to connect the energy and position signals in detector zone 2j to the common output bus 550. The narrow logic high pulse from LC_(2j)525 also appears at the OR_(2j) 545 output. Since it is possible for Vin detector zone 2 j to be also in coincidence with an event W indetector 1, in addition to event U, or with events in other detectorswithin the FOV of detector 2, the block LC_(1i) 520 is used to ensurethat a valid coincidence is limited to just event U in detector zone 1i. The gate AND_(ij) 560 has a logic high output only for a two-waycoincidence between V and U in detectors 2 and 1, thus ruling out athree-way coincidence with other events. When this is the case, andsuitably fast scintillators are being employed, the coincidencevalidation circuitry can also be used to generate an output pulseΔTF_(ij) with duration τ_(c)+ΔTF_(ij), where ΔTF_(ij) is the differencebetween the flight times of gamma rays 1 i and 2 j. TFP_(ij) hasduration τ_(c) if ΔTF_(ij)=0 and duration 2τ_(c) if ΔTF_(ij)=τ_(c). Thenarrow logic-high pulse from AND_(ij) 560 is used to generate a pulse ofwidth T_(p) 570. The pulse ΔTF_(ij) 580 is at logic level 0 when thereis no valid coincidence. When there is coincidence and ΔTF_(ij) is atlogic 1, the logic level of the ΔTS_(ij) 590 shows whether or not gammaray 1 i leads (ΔTS_(ij)=1) or lags (ΔTS_(ij)=0) gamma ray 2 j. Theenergy and position signals of event V in detector zone 2 j arepresented as outputs E_(2j), x_(2j) and y_(2j), respectively, while thepulse of width T_(p) replaces Z_(2j) at the output, as shown in thefigure for M=1. When operation is in SPECT mode, or M=0, the signalsZ_(2j), E_(2j), x_(2j) and y_(2j) in detector zone 2 j are passed on tothe outputs without modification. While FIG. 5 shows only two blocks 1 iand 2 j, it should be kept in mind that i and j are variables (1≦i≦F,1≦j≦F) as noted earlier, and, in fact, F blocks on either side are beingrepresented. When operation is in PET mode, the four lower switches inthe figure ensure that events in the detector 2 are arranged in the sameorder as related coincidence events in detector 1, enabling coincidentsignals in the two detectors to be applied to corresponding ADC sets inthe opposing detectors. The digitized values in the data latches can beread concurrently through the two F-input multiplexers, and thenconsecutively by the two-input multiplexer, as indicated in FIG. 6below. The x- and y-coordinates of coincident events may even be readinto computer memory as single words to facilitate data backprojectionand subsequent image reconstruction.

FIG. 6 illustrates a simplified functional block diagram of an exemplaryhexagonal PET/SPECT scanner in accordance with an embodiment of thepresent invention. The scanner operation is in SPECT mode if M=0 ischosen and in PET mode if the coincidence validation circuitry (CVC) inFIG. 5 is activated by choosing M=1. The coincidence validationcircuitries CVC #1 through CVC #F 625, where F is as defined in above,include detector-zone circuitry up to and including the PHA 280 in FIG.2B. In PET mode, detector D can have coincidences with the detectorhalves B2, A1, A2 and F1 in FIG. 4B. A detector zone along the boundaryof the detector halves D1 and D2 can have valid coincidences only withthe detector halves A1 and A2. A detector zone at the far left ofdetector D1 can have valid coincidences with the detector halves A2 andF1, whereas a detector zone at the far right of D2 can have validcoincidences with the detector halves B2 and A1. In general a detectorzone anywhere in the scanner can have valid coincidences with theequivalent of at most two detector halves, which means with F detectorzones. Coincidence validation circuitry (CVC) 630 is provided to ensurethat only coincidences with these detector zones are accepted and thatcoincidences from outside of the FOV are discarded to minimize scatterand accidental coincidences. TOF capability can be added to the PETscanner provided that a suitably fast scintillator replaces the NaI(Tl).

FIG. 7A illustrates exemplary geometric sampling employing a largernumber of reduced effective size PMTs by using fiber-optic coupling inaccordance with an embodiment of the present invention. Each of theregular PMTs P₁-P₄ in the detector zone is replaced by four PMTs t₁-t₄,or a four-channel PMT of reduced effective size by using fiber-opticcoupling. FIG. 7B illustrates an exemplary equivalent circuit of PMT P₁in FIG. 7A and its PA in accordance with an embodiment of the presentinvention. FIG. 7C illustrates an exemplary equivalent circuit of afour-channel PMT, and associated PAs and SA in accordance with anembodiment of the present invention. Dividing the scintillation lightapplied to one PMT equally between two or more PMTs also reducesnonlinear distortion due to P₁ and the PA, since the four-channel PMTand the PAs now operate in their more linear ranges. Although theoptical fibers are shown to have rectangular terminations in FIG. 7A,the terminations are ideally hexagonal on the scintillator side, and ofthe same shape as the PMTs on the PMT side.

FIG. 8 illustrates an exemplary circuit for four radiation counters, sayt₁ through t₄, also arranged as in FIG. 7A, in accordance with anembodiment of the present invention. Outputs v₁ through v₄ of thecounters share an amplifier 810, a PHA 820, and an ADC 830. A logic highZ pulse is generated each time a valid particle or photon is detected.The switch positions shown are for the case when a valid signal v₁ ispresent. Coincidence detection leading to more than one valid signal ata time is prevented. Optionally, the circuit can include a provision togenerate a logic-high pulse R as shown each time two or more coincidentevents are being rejected. The count of these pulses can be used toprovide statistical information regarding the margin of error with whichthe total count is recorded.

Having fully described at least one embodiment of the present invention,other equivalent or alternative methods and an apparatus for highresolution/sensitivity and improved image signal-to-noise ratio (SNR)for planar and tomographic (PET/SPECT) imaging according to the presentinvention will be apparent to those skilled in the art. The inventionhas been described above by way of illustration, and the specificembodiments disclosed are not intended to limit the invention to theparticular forms disclosed. For example, the particular implementationof the scintillation detectors and photodetectors may vary dependingupon the particular type nuclear imaging scanner. The circuits describedin the foregoing are exemplary and other designs in accordance with theteachings provided herein are contemplated as within the scope of thepresent invention. Furthermore, the techniques described in theforegoing were directed to nuclear imaging; however, similar techniquesmay be applied to various other types of scanners, such as inastrophysics, high energy radiation detectors in nuclear physics, etc.,and are contemplated as within the scope of the present invention. Theinvention is thus to cover all modifications, equivalents, andalternatives falling within the spirit and scope of the followingclaims.

Claim elements and steps herein have been numbered and/or letteredsolely as an aid in readability and understanding. As such, thenumbering and lettering in itself is not intended to and should not betaken to indicate the ordering of elements and/or steps in the claims.

What is claimed is:
 1. A method comprising: steps for choosing theoptimum detector for producing bursts of optical photons in response toevents comprising incident radiation; steps for arranging photodetectorsin a honeycomb array into a plurality of detector zones for concurrentlyconverting said bursts of optical photons into signal outputs therebyachieving high sensitivity; steps for identifying valid events whilerejecting signal outputs due to misplaced pileup events arising fromscattered photons of said incident radiation steps for exchanging excessdetector sensitivity for spatial resolution by significantly reducingthe detector's thickness, which leads to reduction indepth-of-interaction (DOI) blurring and minimizes blurring due toCompton-scattered radiation, enabling nuclear images of unprecedentedresolution and signal-to-noise ratio (SNR) to be obtained; steps forsimultaneously generating energy and position signals for events atmultiple detector zones during a single deadtime period, therebyreducing effective detector deadtime and increasing count-ratecapability; and steps for transferring said energy and position signalsto computer memory by means of a high-speed multiplexer for dataanalysis and image display.
 2. The method as recited in claim 1, furthercomprising steps for mitigating errors arising from inequalities ingains of said photodetectors and associated preamplifiers.
 3. The methodas recited in claim 1, further comprising steps for increasingphotodetector array size and reducing detector deadtime by reducingeffective size of said photo detectors.
 4. A method comprising steps of:detecting energy for producing bursts of optical photons in response toevents comprising incident radiation; arranging F number of sharinggroups of seven photodetectors in a honeycomb array for viewing zones ofsaid bursts of optical photons and converting said bursts of opticalphotons into signal outputs, where each of said central groups isassociated with a detector zone; rejecting signal outputs that are dueto scattered photons of said incident radiation for each of said centralgroups; generating energy and position signals in detection of F numberof said events during a single deadtime period for each of said F numberof sharing groups; and transferring said energy and position signals ofsaid F number of said events to computer memory for image display anddata analysis.
 5. The method as recited in claim 4, further comprisingthe step of reducing effective detector deadtime using pileup preventioncircuitry.
 6. The method as recited in claim 4, further comprising thestep of converting analog energy and position signals to digital.
 7. Themethod as recited in claim 4, further comprising the step of mitigatingdistortions arising from device nonlinearities using spatial linearitycorrection circuitry.
 8. The method as recited in claim 4, furthercomprising the step of mitigating errors arising from inequalities ingains of said photodetectors and associated preamplifiers using energycorrection circuitry.
 9. The method as described in claim 4, furthercomprising the step of coupling said bursts of photons to saidphotodetectors using fiber-optics for reducing effective photodetectorsize for improving detector spatial sampling.
 10. A system comprising:at least one continuous energy detector for producing bursts of opticalphotons in response to events comprising incident radiation; amulti-zone detector architecture comprising F number of sharing groupsof seven photodetectors arranged in a honeycomb array for viewing zonesof said at least one energy detector and converting said bursts ofoptical photons into signal outputs, where each of said central groupsis associated with a zone of said at least one energy detector; amisplaced pile-up suppression (MPS) circuitry associated with each ofsaid central groups for rejecting signal outputs that are due toscattered photons of said incident radiation; a detector-zone circuitryassociated with each of said F number of sharing groups for generatingenergy and position signals in detection of up to F number of saidevents during one deadtime period; pileup prevention circuitry (PPC) forrejecting events in said detector zone following within one deadtimeperiod of a first event; a high-speed multiplexer with F inputs fortransferring said energy and position signals of said F number of saidevents to computer memory for image display and data analysis; anddigital circuitry for the realtime determination of time-of-flight TOF(or TF) information in PET to speed up image reconstruction.
 11. Thesystem as recited in claim 10, further comprising pileup preventioncircuitry (PPC) for reducing effective deadtime and greatly extendingthe count-rate capability of said at least one energy detector.
 12. Thesystem as recited in claim 10, further comprising analog-to-digitalconverters for said energy and position signals.
 13. The system asrecited in claim 10, further comprising spatial linearity correctioncircuitry for mitigating distortions arising from device nonlinearities.14. The system as recited in claim 10, further comprising energycorrection circuitry for mitigating errors arising from inequalities ingains of said photodetectors and preamplifiers.
 15. The system asrecited in claim 10, further comprising fiber-optic coupling betweensaid at least one energy detector and said photodetectors for reducingan effective size of said photodetectors.
 16. The system as recited inclaim 15, wherein said fiber-optic coupling at said at least one energydetector comprises a hexagonal shape for maximizing the packing fractionof said photodetectors.
 17. The system as recited in claim 10, whereinsaid F number of central groups in each of said F number of sharinggroups share said associated detector-zone circuitry to minimizedetector cost.
 18. The system as recited in claim 10, wherein realtimeTOF (or TF) capability using coincidence validation logic circuitry formore efficient use of imaging time and computer memory in place of theexisting method of acquisition of raw data with time stamps forsubsequent processing using software.
 19. The system as recited in claim17, wherein two to seven adjacent radiation or particle counters,arranged hexagonally in the form of a detector zone, can share analogand digital circuitry to reduce the cost while maintaining the spatialresolution capabilities of the individual counter.